Stably and nondestructively positioning artificial materials across living cell membranes is a long-held goal of biomaterials. Accessing the intracellular space of a cell is of essential importance in many different biomedical applications. The ability to specifically and nondestructively incorporate inorganic structures into or through biological membranes is a key step toward realizing full bioinorganic integration in applications such as drug delivery, electrical recording and stimulation, targeted cellular therapeutics, and biosensors. Interfaces for the delivery of inorganic objects across cell membranes generally rely upon destructive formation of membrane holes and serendipitous adhesion, rather than selective penetration and attachment into the bilayer itself. A more benign means to penetrate through the cell membrane is required. While surface modification techniques have been highly successful at controlling cell mobility, proliferation, and differentiation on two-dimensional surfaces, bridging across the cell membrane itself has been much less studied.
In nature, membrane proteins are examples of systems whose outer surface is designed to specifically interact with the interior of the cell membrane lipid bilayer. The tight junction between the lipid and protein eliminates constitutive ion or protein leakage, allowing membrane proteins to regulate the chemical flux through the bilayer. This relationship has been difficult to achieve with man-made biomaterials. Although many man-made biomaterials are designed to regulate the interactions between artificial and natural surfaces, including across the cell membrane, when such materials are inserted through the cell membrane the interface formed between the interior edge of the membrane and the material surface is not well understood and poorly controlled.
While many gene and drug carrier particles appear to enter the cell by endocytotic mechanisms, materials such as cationic polymers and nanoparticles have been shown to directly penetrate the membrane. However, these highly charged species can create holes leading to significant cytotoxicity, and the underlying lipid-cation interaction mechanism is still poorly understood. New materials delivery systems such as DNA functionalized nanowires pierce the membrane and have had some success delivering cargo, but cells are unable to survive longer than several days following penetration.
Thus, nanomaterials and nanostructured surfaces offer new opportunities to interact with biological species at their native length scales, promising more effective interfaces if the appropriate architectures can be discovered. In particular, engineered interfaces between nanostructures and lipid bilayers, themselves nanoscale two-dimensional fluids approximately 5 nm thick, may provide a unique means to breach this defensive wall enabling direct chemical and electrical access to the cell's interior. Technologies for drug delivery, electrical (e.g., ion-channel) measurements, single cell analysis, and gene therapy would all benefit from an improved understanding of how to establish direct chemical and electrical conduits to the cell's interior without inducing detrimental side effects. While there are existing methods for gaining intracellular access, the techniques tend to be destructive (electroporation and patch-clamping), slow (microinjection and patch-clamping), or inefficient (liposomal delivery and endocytotic uptake).
For example, direct electrical access into the cell interior is required for low-noise recording of ion channel activity; yet conventional patch clamp techniques are destructive, leading to rapid cell death, while on-chip devices have poor seal resistances. Yet there is a huge potential benefit if electrodes that nondestructively incorporate into biological membranes could be realized. For example, the patch-clamp technique has been the gold standard for fundamental studies relating to the electrical properties of cells. These experiments have included whole cell behavior down to individual ion channel activity, and have become a critical tool for the discovery of drugs that affect these proteins. The conventional “whole-cell” patch-clamp involves gently pressing a fine glass micropipette 1-2 μm in diameter against the cell membrane and applying suction, tearing a hole in the membrane for intracellular access, and forming a high-resistance seal with the membrane. While the actual structure of the membrane-pipette interface is not understood, the patch-clamp technique is highly successful at forming tight seals with resistances of several gigaohms. This gigaohm seal enables the measurement of ion channel currents with extremely high signal-to noise ratio. However, the rapid apoptosis of patched cells (approximately 2 h), small number of simultaneous measurements (e.g., 2-3), and inherently serial process have limited applicability of conventional patch-clamping for monitoring cell behavior over extended periods. To overcome these limitations, planar chip-based and automated patch clamp devices have been developed. These existing chip-based electrodes are based on arrays of micromachined holes formed in the substrate or suspended insulating layer as on-chip facsimiles of the pipette tip. Although successful at accelerating the patch clamping process, these devices suffer from low gigaohm seal formation rates and the same rapid cell apoptosis as traditional patch-clamp methods. Efforts to improve seal resistance have largely focused on homogeneously modifying the inorganic materials, including silicon oxide coated nitride membranes, silicon coated with plasma-enhanced chemical-vapor deposited oxide, silicon elastomers, polydimethylsiloxane, glass and quartz and varying the surface roughness, all with limited success. Ideally, chip-based solid state electrodes could enter the cell without causing cell apoptosis, yet previous attempts could not achieve the high-resistance membrane seals necessary.
The lipid bilayer itself is composed of two lipid leaflets, and consists of three different zones: external hydrophilic head groups, hydrophobic lipid tail groups that form the lipid core, and internal hydrophilic head groups. In their liquid state, lipids in the bilayer are highly mobile species, with the structure a dynamic balance between the hydrophobicity of the tail groups, stretching/compression of the tails, head group repulsion, and the relative head to tail dimensions. Because of this mobility, a wide array of lipid bilayer structures has been observed, including vesicles, lamellar sheets, triple junctions, tubules, and platelets. Small molecules and even some nanoparticles have been shown to be able to partition into and through the bilayer without dramatically altering its organization. However, when larger materials penetrate the bilayer an edge or interface must be created. The structure of this new interface is not clear, yet would be anticipated to depend upon the material's nanoscale morphology and hydrophobicity.
FIG. 1 shows four scenarios for bilayer interface structure after material penetration. The first (FIG. 1A) is an idealized ‘fused’ state, where the bilayer makes intimate contact with the probe with little or no disruption of the lipid organization. This is most reminiscent of transmembrane protein interfaces, which often have the first layer of lipids transiently adsorbed on the protein surface. The uninterrupted hydrophobic layer and tight interface serve as a significant barrier for ion or fluid flow, preventing exchange from one side of the membrane to the other. Indeed, cell membranes can have electrical resistances ranging from 10-100 Gigaohms (GΩ) implying almost no ion leakage occurs at the thousands of protein-membrane interfaces. This scenario likely requires nanoscale modification of the probe surface, since it simultaneously interacts with both the hydrophilic and hydrophobic zones of the lipid bilayer.
The second structure (FIG. 1B) is the ‘T-junction’ configuration. This architecture essentially splits the bilayer into two monolayers in which each contact the probe surface, similar to triple-bilayer junctions observed in the hemifusion state during membrane fusion. This arrangement may be energetically favorable for hydrophobic probes, since the surface is in contact with the hydrophobic bilayer tails. A key aspect to this arrangement is the formation of an unfavorable empty interstice, or void, where the bilayer splits, estimated to cost about 10 kbT per nm length of interstice. In hemifused junctions between flexible lipid vesicles this energy can be reduced by increasing the local curvature and lipid splay, however in this case the bilayer must conform to the stiff probe surface and is thus largely unable to do so. This state may therefore be weaker than the fused state. Related structures, such as membrane stalks, have been predicted to increase the rate of hole formation in the surrounding membrane, which could also destabilize this interface. The third situation (FIG. 1C) is the ‘ruptured’ state where the bilayer forms a hole around the probe with a hydrophilic lipid edge near the probe surface but not in direct contact. This may be the favored configuration for hydrophilic probe surfaces since both materials are in continuous contact with water, however, the gap allows fluid and ions to diffuse through the interface. The leakage rate through the junction could vary greatly depending upon the separation between the edge and the probe surface and may fluctuate over time. Energy considerations and molecular dynamics simulations imply that the ruptured bilayer edge consists of a hemispherical cap of lipids which shield the hydrophobic core from the aqueous phase. The curvature and partial exposure of the tail groups make the edge a relatively high energy state, with line energies on the order of 10 pJ m−1. This is still considerably smaller than direct lipid tail-water contact, which can be estimated using typical alkane-water surface energies of 25 mJ m−2 to be roughly 75 pJ m−1 for a 3 nm thick bilayer core. While at equilibrium holes in the membrane are unusual, they are commonly created by artificial means such as electroporation or mechanical tension, thus are not unlikely for penetrating probes.
The final scenario (FIG. 1D) is the ‘adhered’ state, in which the lipid surface is attached to the surface of the probe. Lipid-surface adhesion is common in supported lipid bilayers, driven by electrostatic and van der Waals attractions. The lipid is usually not in direct contact with the surface, instead separated by a 1-2 nm aqueous gap. This gap allows some ion transport through the junction as measurements have found conductivities of approximately 0.002 Ω−1 cm−1 for model lipid/glass interfaces, yet larger proteins may be prevented from passing. This is thought to be the interface created during patch-clamp measurements, and with sufficient surface contact area could be highly resistive.
The specific nature of the probe/lipid junction is important for the mechanical strength, electrical resistivity, and cytotoxicity of the interface. For non-destructive interfaces to cells and membranes, the fused state is likely optimal by preventing leakage from the cytosol and maintaining strong attachment. However, to date, a fused seal such as shown in FIG. 1A has not been possible to reliably achieve. Described herein are devices, systems, and methods for forming stable, long-lasting and non-destructive fused seals with cells. The devices, systems and methods described herein may address the goals and limitations of other technologies discussed above.
These interfaces may also be crucial for interfacing to neurons located in the brain tissue. Within the brain, interconnected networks of cells underlie all cognitive functions, yet monitoring the dynamic signaling networks at the single cell level have not been possible. In particular, Local Field Potential (LFP) recordings only provide ensemble activity information, burying the individual cell contributions and making it impossible to pull out local cell function within the network. In theory, this limitation might be overcome by both shrinking the size of the electrode to match a single cell (˜10 um), and deployed at high enough density to capture a sufficient percentage of cells within a circuit. However, there are significant challenges to this approach. Even at cellular size scales and sensitivity, there is no guarantee that the electrode will pick up signal only from a single cell as the LFP is usually an order of magnitude larger. Secondly, micromotions are problematic. Micromotions are problematic even for larger-scale devices, but as the size scale of the sensor is reduced even more, nanoscale changes in the proximity between the sensor and the cell may drastically alter the recorded signal. Thirdly, since the recordings would come from cells in the immediate vicinity of the electrodes, inflammation and damage from large electrode substrates can prevent single cell pickup. Finally, the neuron that the electrode is actually measuring electronically is typically not visible by other means, such as optical microscopy, limiting the integration with other techniques such as optogenetics. Thus, while conventional microelectrodes could be scaled to single cell size and high density, these additional concerns make this simplistic approach impractical.
Currently several different electronic recordings platforms exist for in-vivo recordings. The canonical devices consist of the Utah Array, a bed-of-nails arrangement of silicon microneedles, and the Michigan Array, a planar array of electrodes on a millimeter-scale spike. Microfabrication technology makes it possible to create large arrays of metal electrodes, consisting of hundreds to thousands of devices, each one of which can easily be scaled to the single cell level (˜10 um). Demonstrations of this type of electrode on flexible polyimide substrates reach back until the early 1990's, and today well over a hundred papers on these devices are available, many with in vivo data. However, such devices have not yet achieved large-scale, single cell recordings, likely because of the way such electrodes interface to the surrounding cells. For planar electrode arrays, there is no interaction between the cells and the electrodes, such that the electrical signal from the nearest cell is entirely dependent on the somewhat random distance between the cell and the electrode produced after insertion. Moreover, the local field potentials (LFP), which is on the order of 1-5 mV for typical Au electrodes, will overwhelm the typical 100-200 uV extracellular potential, even when the cells are in direct contact with the electrode as seen in multi-electrode arrays (MEAs). Scaling the electrodes to neuronal dimensions (˜10 um for a single cell) and smaller than the inter-neuronal distance (˜25 um) also make these electrodes particularly prone to noise and nanoscale changes in the distance to the nearest cell. Thus, a method to both couple a single cell to one electrode and screen out LFP environmental signals is desirable.